U.S. patent number 6,285,740 [Application Number 09/419,261] was granted by the patent office on 2001-09-04 for dual energy x-ray densitometry apparatus and method using single x-ray pulse.
This patent grant is currently assigned to The United States of America as represented by the Secretary of the Navy. Invention is credited to Craig N. Boyer, Glenn E. Holland, John F. Seely.
United States Patent |
6,285,740 |
Seely , et al. |
September 4, 2001 |
Dual energy x-ray densitometry apparatus and method using single
x-ray pulse
Abstract
A dual-energy x-ray system is provided which is useful in
densitometry and other applications. A dual energy x-ray source,
formed field emission x-ray tube driven by a Marx generator,
produces a single x-ray pulse of a very short duration (e.g., 10 to
200 nanoseconds). The pulse provides x-ray energy of a first, high
value early in the pulse and x-ray energy of a second, lower value
later in the pulse so as to provide a dual energy level x-ray
distribution comprising hard and soft x-rays. A detector system
having dual x-ray energy discrimination properties, formed by soft
and hard x-ray detectors and an inter-detector, receives the x-ray
pulse and discriminates between the dual x-ray energy levels of the
pulse.
Inventors: |
Seely; John F. (Fairfax,
VA), Boyer; Craig N. (Mitchellville, MD), Holland; Glenn
E. (Wheaton, MD) |
Assignee: |
The United States of America as
represented by the Secretary of the Navy (Washington,
DC)
|
Family
ID: |
23661493 |
Appl.
No.: |
09/419,261 |
Filed: |
October 13, 1999 |
Current U.S.
Class: |
378/98.9;
250/367; 250/370.09; 250/370.11; 378/101; 378/102; 378/106 |
Current CPC
Class: |
H05G
1/20 (20130101); H05G 1/30 (20130101) |
Current International
Class: |
H05G
1/20 (20060101); H05G 1/00 (20060101); H05G
1/30 (20060101); H05G 001/64 () |
Field of
Search: |
;378/98.9,98.11,101,102,103,106,122 ;250/367,370.11,370.09 |
References Cited
[Referenced By]
U.S. Patent Documents
Other References
JF. Seely et al., Dual-Energy Bone Densitometry Using a Single 100
ns X-Ray Pulse, 25 Med. Phys. 2027 (No. 10, Oct. 15,
1998)..
|
Primary Examiner: Kim; Robert H.
Assistant Examiner: Ho; Allen C.
Attorney, Agent or Firm: Karasek; John J. Ferrett; Sally
A.
Claims
What is claimed is:
1. A dual-energy x-ray system comprising:
a dual energy x-ray source, comprising a field emission x-ray tube
driven by a Marx generator, for producing a single x-ray pulse
providing x-ray energy of a first value early in the pulse and
x-ray energy of a second, lower value later in the pulse so as to
provide a dual energy level x-ray distribution comprising hard and
soft x-rays; and
a detector system, having dual x-ray energy discrimination
properties, for receiving said pulse and for discriminating between
the dual x-ray energy levels of said pulse.
2. A dual-energy x-ray system as claimed in claim 1 wherein said
detector system comprises, in tandem, a soft x-ray detector, a hard
x-ray detector and an inter-detector filter disposed between said
soft and hard x-ray detectors for filtering soft x-rays passing
through said soft x-ray detector.
3. A dual-energy x-ray system as claimed in claim 2 further
comprising a pre-filter at the output of the x-ray tube for
filtering x-rays produced by said x-ray tube to enhance the dual
level x-ray energy distribution, harden x-ray fluence and reduce
x-ray dosage to an object receiving said x-ray pulse.
4. A dual-energy x-ray system as claimed in claim 2 wherein said
soft x-ray detector comprises a phosphor mounted on an electronic
detector panel and said hard x-ray detector comprising further
phosphor mounted on a further electronic detector panel arranged in
tandem with said first panel.
5. A dual-energy x-ray system as claimed in claim 2 wherein said
soft x-ray detector comprises a phosphor and said hard x-ray
detector comprises a scintillator.
6. A dual-energy x-ray system as claimed in claim 5 wherein said
detector system comprises a first fiber optic plate, wherein said
phosphor is supported by said plate, and wherein said scintillator
comprises a scintillator element supported by a further fiber optic
plate and forming an integral structure with said first plate.
7. A dual-energy x-ray system as claimed in claim 5 wherein said
phosphor and said scintillator produce photons of different colors
and said detector system further comprises an imaging camera for
producing different colored images and CCD detectors for separately
capturing said different colored images.
8. A dual-energy x-ray system as claimed in claim 1 wherein said
pulse produced by said x-ray source has a duration of between 10
and 200 nanoseconds.
9. A dual-energy x-ray system as claimed in claim 8 wherein said
pulse has a duration of about 100 nanoseconds.
10. A dual-energy x-ray system as claimed in claim 8 wherein said
pulse has a voltage of 30-300 keV.
11. A method of determining bone mineral density, said method
comprising:
using an energy x-ray source comprising a Marx generator and a
field emission x-ray tube driven by the Marx generator to produce a
single x-ray pulse of a duration between 10 and 200 nanoseconds
having two different x-ray energy levels during the duration of the
pulse so as to provide an x-ray distribution within the pulse
comprising both soft and hard x-rays, and
separately detecting the two different x-ray energy levels.
12. A method as claimed in claim 11 wherein the two different x-ray
energy levels are detected using a detector system comprising a
soft x-ray detector and a hard x-ray detector and wherein an
inter-detector filter is disposed between said soft and hard x-ray
detectors for filtering soft x-rays passing through said soft x-ray
detector.
13. A method as claimed in claim 12 further comprising providing a
pre-filter for filtering x-rays produced by said x-ray tube to
enhance the dual level x-ray energy distribution, harden x-ray
fluence and reduce x-ray dosage to an object receiving said x-ray
pulse.
14. A method as claimed in claim 12 wherein the soft x-ray detector
used comprises a phosphor and the hard x-ray detector used
comprises a scintillator, wherein said phosphor and said
scintillator produce photons of different colors, wherein different
colored images are produced using an imaging camera and wherein
colored images are separately captured using the different CCD
detectors.
15. A method as claimed in claim 11 wherein the pulse produced has
a voltage of 30-300 keV.
Description
FIELD OF THE INVENTION
The present invention relates to x-ray systems for determining bone
mineral density and for other applications, and to methods using
such systems.
RELATED ART
Precision bone mineral densitometry is important for the early
detection of osteoporosis and the prediction of future bone
fracture risk. Bone mineral loss is associated with aging and is
more rapid in post-menopausal women. In addition, bone mineral loss
is accelerated during long-term bed rest and in the weightless
environment of space.
Prior clinical studies indicate that the association of measured
bone density with osteoporosis and bone fracture is more
significant for the major weight-bearing axial sites (lumbar spine
and proximal femur) than for extremity sites (hand, radius, and
calcareous). The bone mineral content is correlated with the
vertebral strength determine in vitro. Clinical studies also
indicate that the preferred site for bone mineral densitometry may
be the lumbar spine, and the preferred view is lateral rather than
anterior-posterior (AP). The lateral view of the vertebrae is not
obscured by the posterior vertebral elements. This permits the
isolation of the trabecular bone region that is most susceptible to
mineral loss. However, soft tissue absorption can affect lateral
projection measurements to a greater extent than AP projection
measurements.
It has been estimated that a measured bone mineral density (BMD)
that is one standard deviation (approximately 2%) below the average
for a control population implies a significantly higher risk of a
future bone fracture. Thus, any bone densitometry technique should
have at least 2% precision and absolute accuracy.
In many clinical settings, it is important to determine BMD with a
short patient observation time. In addition, since BMD measurements
may be repeated over a period of time, the exposure to the patient
should be as low as possible for each observation. Although
ultrasonic and magnetic resonance methods are convenient and have
no x-ray exposure, such methods do not have the required accuracy
at the present time.
The rather high precision and accuracy required for bone
densitometry has resulted in the development of x-ray radiographic
absorption techniques. Bone densitometry apparatus for the hand,
radius, and calcareous have been developed. However, these
measurements are not as significant as axial site measurements for
diagnosing osteoporosis and predicting fracture risk.
Considering some of the techniques that have been used,
quantitative computed tomography (QCT) x-ray scanning techniques
produce three-dimensional images of skeletal regions. This permits
the elimination of soft tissue attenuation and a precise
determination of the BMD of interior trabecular bone regions.
However, QCT techniques are not suitable for repeated clinical
measurements because of the high patient dose (surface dose
.about.100 mrem), the long patient positioning and immobilization
time (tens of minutes), and the relatively high cost of the
apparatus.
Dual-energy x-ray absorptiometry (DEXA) projection scanning units
have been developed by Hologic Inc. (Waltham, Mass.) and Lunar
Corp. (Madison, Wis.). Using dual-energy calibration and
subtraction algorithms, the attenuation by soft tissue is
effectively eliminated from the measurement, and the BMD of axial
sites can be precisely and accurately determined. In the case of
lumbar spine and proximal femur measurements, clinical studies
indicate a typical precision of 1-3% and an absolute accuracy of
4-15%. The patient surface dose is typically of order 10 mrem while
the patient positioning and observation time is approximately 10
minutes.
Lumbar vertebrae have been characterized in vitro by ash mass,
metrology, and bone densitometry techniques. The directly measured
quantities have been compared to the equivalent quantities
determined in vitro by non-invasive techniques. The measured ash
mass (mineral content excluding water, soft tissue, and other
combustibles) is in the range 5-15 g. The ash mass is equivalent to
the bone mineral content (BMC). The volume measured by metrology is
in the range 30-50 cm.sup.3, and this quantity may be determined by
QCT techniques. The projected area measured by metrology is in the
range 11-17 cm.sup.2, depending on the viewing direction, and can
be determined by radiographic projection techniques. The ash mass
to volume values are in the range 0.20-0.35 g/cm.sup.3, and this
quantity is equivalent to the bone mineral volumetric density
determined by QCT. The ash mass to projected area values are in the
range 0.4-1.2 g/cm.sup.2. This quantity is equivalent to the bone
mineral areal density that is determined by most DEXA instruments,
and this quantity is usually referred to as the bone mineral
density (BMD).
The primary goal for a DEXA instrument is to determine the BMD to a
precision of 2%. In addition, the absolute BMD, typically in the
range 0.4-1.2 g/cm.sup.2, should be determined to an accuracy of
2%. The primary difficulty is to account for the soft tissue areal
density, along the same line of sight, with values up to 20 to 30
g/cm.sup.2 in the case of the lumbar spine.
In general, a dual-energy x-ray distribution useful in BMD
measurements may be enhanced by switching the source voltage
(developed by Hologic Inc.), changing the filtration (developed by
Lunar Corp.), or a combination of these two techniques. Previous
studies that were based on x-ray tube loading, x-ray quantum noise,
and patient exposure established that x-ray energies in the two
ranges 40-60 keV and 80-130 keV are suitable for dual-energy x-ray
absorptiometry. Most DEXA systems are limited by x-ray quantum
noise, and other noise sources (such as phosphor, detector, and
electronics) are smaller. As indicated above, patient exposure is
of primary concern, although the exposure provided by DEXA systems
is usually lower than the exposure from other radiology procedures
and from the natural background.
SUMMARY OF THE INVENTION
In accordance with the invention, a system is provided which
employs a pulsed, portable hard and soft x-ray source which is
useful in medical imaging in general, and BMD measurements in
particular, as well as in flash x-ray absorptiometry. As discussed
below, the x-ray source produces a single very short (10 to 200
nanosecond) pulse, and this is of obvious advantage in treating
patients, particularly as compared with systems such as those
described above which involve substantially longer patient exposure
times. Further, the loading on the x-ray tube is minimal during the
single x-ray pulse, thereby permitting operation of the x-ray tube
at higher voltage and power levels than is possible with
conventional x-ray tubes. This is important in patient diagnosis
because a higher x-ray number density reaches the associated
detector, thereby improving the x-ray quantum signal to noise ratio
and reducing the dose absorbed by the patient.
According to one aspect of the invention, a dual-energy x-ray
system is provided which comprises: a dual energy x-ray source,
comprising a Marx generator and a field emission x-ray tube driven
by the Marx generator, for producing a single x-ray pulse providing
x-ray energy of a first value early in the pulse and x-ray energy
of a second, lower value later in the pulse so as to provide a dual
energy level x-ray distribution comprising hard and soft x-rays;
and a detector system, having dual x-ray energy discrimination
properties, for receiving said pulse and for discriminating between
the dual x-ray energy levels of the pulse.
Preferably, the detector system comprises, in tandem, a soft x-ray
detector, a hard x-ray detector and an inter-detector filter
disposed between the soft and hard x-ray detectors for filtering
soft x-rays passing through the soft x-ray detector.
A pre-filter at the x-ray tube output is preferably provided for
filtering x-rays produced by the x-ray tube to enhance the dual
level x-ray energy distribution, harden x-ray fluence and reduce
the x-ray dosage to an object receiving the x-ray pulse.
In one preferred embodiment, the soft x-ray detector comprises a
phosphor mounted on an electronic detector panel and the hard x-ray
detector comprising a further phosphor mounted on a further
electronic detector panel arranged in tandem with the first
panel.
In an advantageous implementation, the soft x-ray detector
comprises a phosphor and the hard x-ray detector comprises a
scintillator.
In a second preferred embodiment employing the aforementioned
implementation, the detector system comprises a first fiber optic
plate, the phosphor is supported by that plate, which is, in turn,
optically coupled to a second scintillating fiber optic plate to
form an integral structure with the first plate. Advantageously,
the phosphor and scintillator produce photons of different colors
and the detector system further comprises an imaging camera for
producing different colored images and CCD detectors for separately
capturing the different colored images.
As indicated above the pulse produced by the x-ray source
advantageously has a duration of between 10 and 200 nanoseconds. In
a typical application, the pulse has a duration of about 100
nanoseconds. In a preferred implementation, the voltage of the
pulse is between about 30-300 keV.
In accordance with a further aspect of the invention, a method of
determining bone mineral density is provided, the method
comprising: using an energy x-ray source comprising a Marx
generator and a field emission x-ray tube driven by the Marx
generator to produce a single x-ray pulse of a duration between 10
and 200 nanoseconds having two different x-ray energy levels during
the duration of the pulse so as to provide an x-ray distribution
within the pulse comprising both soft and hard x-rays, and
separately detecting the two different x-ray energy levels.
Further features and advantages of the present invention will be
set forth in, or apparent from, the detailed description of
preferred embodiments thereof which follows.
BRIEF DESCRIPTION OF THE DRAWINGS
FIG. 1 is a schematic block diagram of a dual energy x-ray system
incorporating a dual panel detection scheme in accordance with a
first preferred embodiment of the invention;
FIG. 2 is a schematic block diagram similar to FIG. 1 of a dual
energy x-ray system incorporating a dual color detection scheme in
accordance with a second preferred embodiment of the invention;
FIGS. 3(a), 3(b) and 3(c) are diagrams of the operating parameters
of the x-ray tube of FIGS. 1 and 2 illustrating, respectively, the
calculated current, voltage and power; and
FIGS. 4(a) and 4(b) are diagrams illustrating, respectively, the
energy delivered to the x-ray tube, and the x-ray tube flux.
DESCRIPTION OF THE PREFERRED EMBODIMENTS
Referring to FIG. 1, there is shown a first embodiment of the
invention. The system of this embodiment referred to a dual energy
x-ray source 10 which comprises a Marx generator 10a that drives a
field emission x-ray tube 10b containing a tungsten anode and a
mesh cathode in a vacuum. The source 10 produces a 30-300 keV x-ray
pulse of 100 nanoseconds (ns) duration. In practice the duration of
the pulse can be between about 10 ns and 200 ns. It will be
appreciated that this duration is very much shorter than that
produced by a conventional x-ray source such as is normally used
for the present purposes. As indicated above and explained in more
detail below the unit 10 produces both soft x-rays 12 and hard
x-rays 14.
A pre-filter 16, which can comprise a 0.1 mm tantalum pre-filter,
hardens the fluence, enhances the dual-energy x-ray distribution
and reduces low energy x-ray fluence. Initial filtering is provided
by the 0.2 thick aluminum vacuum window (not shown) of the x-ray
tube 10b of unit 10.
The tissue and bone of a patient on which a bone mineral density
(BMD) measurement is to be made are indicated, in FIG. 1, at T and
B. After passing through the tissue T and bone B the soft x-rays 12
are received by a soft x-ray phosphor and detector 18, i.e. a front
panel or cassette that is sensitive to the soft x-ray component,
and the soft x-ray image is extracted therefrom, while the hard
x-rays 14 pass through detector 18 as well as an inter-detector
filter 20 to a hard x-ray phosphor and detector, i.e., a rear panel
or cassette 22 that is sensitive to the hard x-ray component. The
detector 18 preferably comprises a front phosphor screen that is
optimized to absorb the soft x-ray component while the detector 22
preferably comprises a rear phosphor screen that absorbs the
remaining hard x-ray component. The inter-detector filter 20
enhances the dual-energy discrimination by preferentially
attenuating the residual soft x-rays that pass through the front
phosphor screen comprising detector 18.
Discharge of Marx generator 10a into the field emission tube 10b
delivers to tube a short, high power, pulse which causes explosive
field emission of electrons at the tube's cathode, resulting in a
plasma which produces a dual energy x-ray. Specifically, the
electrostatic field across the tube's anode and cathode extracts
electrons from the plasma, and accelerates them such that they
impact the anode to produce characteristic and Bremstrahlung
x-rays. As Marx generator 10a discharges, the intensity of the
discharge diminishes, resulting in higher energy x-rays at the
outset, and lower energy x-rays at the end of the pulse, thus
producing a dual energy x-ray pulse. Because the generator's pulse
is short and sharp, the x-ray pulse is of short duration but high
intensity. Member 18 is selected to detect the low energy portion
of the pulse, and member 22 the high energy portion, with filter 20
disposed between them to absorb whatever low-energy portion of the
pulse traversed member 18. It is known that one can make Marx
generators small and compact. See, U.S. patent application Ser.
Nos. 09/215,499 and 09/162,150, now U.S. Pat. Nos. 6,166,459 and
6,064,718, the disclosure of each being incorporated herein by
reference. By addition of a dual energy detector such as 18, 20,
22, one has a dual energy x-ray system that is small, compact, and
effective for applications such as bone density scans.
Referring to FIG. 2, there is shown a second, dual-color detection,
embodiment of the invention. The embodiment of FIG. 2 is similar to
that of FIG. 1, and like elements have been given the same
reference numbers. FIG. 2 differs from FIG. 1 with respect to the
detection provided. The detection arrangement of FIG. 2 includes a
soft x-ray phosphor 24, an inter-detector filter fiber optic
element or wafer 26 and a hard x-ray scintillation fiber optic
element or wafer 28. Different colors are emitted by the phosphor
24 and scintillator 28 and a camera (prism) 30 transmits the two
colors to two different CCDs 34 and 32, which produce the
respective soft and hard x-ray images 38 and 36, respectively. This
embodiment is particularly useful, for example, with small areas of
x-ray coverage in contrast to the relative large and quite
expensive detector panels 18 and 22 of FIG. 1.
Calculated values for the x-ray tube current, voltage, and power
are shown in FIGS. 3(a), 3(b) and 3(c) respectively. The initial
electron acceleration phase produced by x-ray source 10 is
characterized by high tube perveance with a rapid increase in the
tube current and voltage. This is followed by an arc phase in which
the current continues to increase while the voltage and impedance
across the tube rapidly collapse in the ensuing vacuum arc. The
x-ray energy is relatively high early in the pulse (when the
electron acceleration voltage is high) and decreases later in the
pulse. The total charge delivered to the anode-cathode gap in the
example under consideration is 0.1 mAs.
The energy delivered to the anode-cathode gap of the x-ray tube
10b, as a function of the electron acceleration voltage, was
calculated from the time-dependent current and voltage. Based on
calculated values, the energy distribution, in units of mJ per kV
of anode-cathode gap voltage, is shown in FIG. 4(a). The energy
delivered at the higher voltages (>150 kV) occurs primarily
during the time of the current spike early in the current pulse
when the voltage is maximal (see FIGS. 3(a) and 3(b)). The energy
delivered at voltages of 30-150 kV occurs primarily during the
middle of the current pulse when the current is high and the
voltage is decreasing. This illustrates the dual-energy nature of
the x-ray source 10, with harder x-rays emitted during the initial
charge-limited electron transport across the anode-cathode gap, and
softer x-rays emitted just prior to the arc phase of the
discharge.
In the example under consideration, the x-ray spectrum was
calculated from the current and voltage using a thick-target
Bremsstrahlung model and accounting for the tungsten K.sub..alpha.
and K.sub..beta. characteristic x-ray lines. The energy in the
characteristic x-ray lines is 4.7 mJ, and the energy in the
continuum is 63 mJ. Of the 10 J energy that is initially sorted in
the energy storage capacitors, 9.8 J is delivered to the x-ray
tube. The efficiency of conversion of the tube energy to x-ray
energy is 0.7%.
With the present invention, in contrast to prior art DEXA systems,
tube loading during a single x-ray pulse is minimal and is not a
constraint on the design of the system. This permits the operation
of the x-ray tube at higher voltage and power levels than is
possible with more traditional x-ray tubes. The higher-energy
x-rays are of diagnostic importance because a higher x-ray number
density reaches the detector, thus improving the x-ray quantum
signal to noise ratio and reducing the patient absorbed dose.
The x-ray detection system of the invention also has
energy-discrimination capabilities. The prior art dual-energy
detection systems discussed above are typically composed of a front
panel or cassette that is sensitive to the soft x-ray component and
a rear panel or cassette that is sensitive to the hard x-ray
component. According to one aspect of the invention discussed
above, energy discrimination is enhanced by positioning the
inter-detector filter (20 or 26) between the two panels to
attenuate the soft x-rays that pass through the front panel.
The hard x-ray images produced by the embodiments of FIGS. 1 and 2
are representative of the skeletal features, while the soft x-ray
image has considerable contribution from soft tissue absorption.
After the soft and hard x-ray images are captured, conventional
image de-composition techniques may be used to subtract the soft
tissue contribution from the hard x-ray image. In addition, the
exposure in the rear panel image by x-rays scattered in the front
panel may be substantially removed.
In most analyses of DEXA systems, the signals are calculated using
the attenuation coefficients of the filters and the object under
study and the sensitivity of the detector. Scattered radiation is
usually not considered for the design of the system. In any case,
scattered radiation can be characterized and accounted for as part
of the in situ calibration process.
Account can be taken for the energy-dependent attenuation of the
x-ray fluence by the materials that are listed in Table 1 below,
for the embodiment of FIG. 2: the aluminum window of the x-ray tube
10b, the tantalum pre-filter 16 that hardens the x-ray fluence and
enhances the dual-energy distribution, the variable tissue areal
density (0-30 g/cm.sup.2), the variable bone mineral areal density
(0-1.5 g/cm.sup.2), the phosphor screen 24 that is designed to
absorb the soft x-ray component, the inter-detector filter 26, and
the scintillation fiber optic 28 that absorbs the hard x-ray
component. The material areal densities, volumetric densities, and
thicknesses are listed in Table 1. For each material, the
energy-dependent attenuation coefficient was derived from the
elemental compositions listed in Table 1. The attenuation
coefficient of tissue was assumed to be that of water. The bone
mineral is hydroxyapatite (Ca.sub.5 P.sub.3 O.sub.13 H) with an
assumed density of 0.25 g/cm.sup.3 based on the lumbar vertebra
metrology. The soft x-ray band, at energies less than the tantalum
K absorption edge at 69.5 keV, contains the tungsten K.sub..alpha.
and K.sub..beta. radiation near 58 keV and 67 keV, respectively,
The hard x-ray band is maximal at 100 keV.
TABLE 1 Material compositions, densities, and thicknesses. Areal
Elemental Density Density Thickness Material Composition
(g/cm.sup.3) (g/cm.sup.2) (mm) Aluminum x-ray Al 2.7 0.0554 0.2
tube window Tantalum Ta 16.6 0.17 0.1 pre-filter Tissue (water)
H.sub.2 O 1.0 0-30 0-300 Bone mineral Ca.sub.5 P.sub.3 O.sub.13 H
0.25 0-1.5 (hydroxyapatite) Soft x-ray phosphor Gd.sub.2 O.sub.2
S:Eu 7.5 0.15 0.2 Barium glass inter- SiO.sub.2 (Ba) 3.37 1.7 5
detector filter Hard x-ray scintillator SiO.sub.2 (Gd.sub.2 O.sub.2
S:Tb) 5.0 7.5 15 Soft tissue C.sub.5 H.sub.8 O.sub.2 1.19 0-30
0-252 phantom (Plexiglas) Bone phantom TiO.sub.2 4.05 0-1.5 0-3.7
(titanium dioxide)
Turning to an example, for the purpose of determining the
coefficient of variation of the BMD measurements, the x-ray fluence
was calculated after passing through variable densities of soft
tissue (0-30 g/cm.sup.2) and bone mineral (0-1.5 g/cm.sup.2). The
attenuation of the soft x-ray component is greater than of the hard
component and the energy distribution is considerably hardened by
the attenuation of the soft x-ray component by bone mineral, which
is important for the sensitivity of the dual-energy technique for
the precise determination of the BMD. In general, the soft x-ray
energy component (40-70 keV) is designed to span the energy range
where the attenuation of bone mineral (-1.0 g/cm.sup.2) is rapidly
decreasing with x-ray energy. The hard x-ray energy component is
designed to penetrate the large soft tissue density (.about.20
g/cm.sup.2).
In this example, based on the embodiment of FIG. 2, the phosphor 24
was chosen to be Gd.sub.2 O.sub.2 S doped with Eu which fluoresces
in the red region of the visible spectrum. The efficiency of
conversion of x-ray energy to visible light energy is 15% (Levy
Hill Laboratories Ltd, Chestnut, United Kingdom). The K absorption
edge of Gd is at 50.2 keV and x-rays with energies above this
value, including the tungsten K radiation near 58 keV and 67 keV,
are efficiently absorbed.
The inter-detector filter 26 advantageously comprises a glass fiber
optic plate with a high barium content (33% by weight; Schott Fiber
Optics Inc., Southbridge, Mass.) and a thickness of, e.g., 5 mm.
The barium K absorption edge is at 37.4 keV, and the tungsten K
radiation is strongly attenuated, as are x-ray energies up to about
100 keV. Because of filter 26, the x-ray fluence that is absorbed
by the rear scintillator 28 is primarily the hard x-ray component
with energies above 100 keV.
The scintillator 28 advantageously comprises a 15 mm thick glass
fiber optic plate doped with Gd.sub.2 O.sub.2 S:Tb (LG-9 plate;
Schott Fiber Optics, Southbridge, Mass.). This material
scintillates in the green region with an efficiency of 15%. The
scintillator 28 absorbs essentially all of the x-ray quanta with
energy below 200 keV.
In considering the x-ray fluences that are absorbed by the aluminum
window of x-ray source 10, the tantalum pre-filter 16, the object
(tissue T and bone mineral B), and the detector components
(phosphor 24, inter-detector filter 26, and scintillator 28), of
primary interest are the fluences absorbed by the front phosphor 24
and the rear scintillator 28. As indicated above, the phosphor 24
preferentially absorbs the soft x-ray component 12 and the
scintillator 28 of the hard component 14 and thus provides the
basis for the dual-energy de-composition of the tissue and bone
mineral attenuation.
For the purpose of determining the BMD, it was assumed that the red
photons from the phosphor 24 (the soft x-ray image) and the green
photons from the scintillator 28 (the hard x-ray image) are
collected and separately detected as provided in FIG. 2 using the
imaging cameras 30 with associated CCD detectors 32 and 34 which
separately capture the red and green images (XC-003 RGB 3CCD
camera, Sony Corp.). The computer model used accounts for the
transmission of red photons through the phosphor 24 and the
channeling of the red photons through the barium glass of the
filter 26 and the scintillation fiber optic plates of
scintillator-detector 28 to the imaging system and camera 30. The
green photons are channeled through the scintillation fiber optic
plate of detector 28 to the imaging system and camera 30. It was
assumed that the collection efficiency of the imaging system is 10%
and that the detection efficiency of the camera system is 1%.
It can be shown that the number of visible light photons that are
generated and detected is much greater than the number of x-ray
quanta that are absorbed. Thus x-ray quantum noise is the dominant
contributor to the coefficient of variation of the BMD.
Before considering the BMD determination process further, it will
be understood that the BMD determining method described here, i.e.,
of separately detecting the optical photons generated by the soft
and hard x-ray components can also use the embodiment of FIG. 1
wherein the two tandem electronic detection panels 18 and 22 are in
contact with the phosphor and the scintillator. In an advantageous
implementation of this embodiment, the electronic detection panels
18 and 20 are of the amorphous silicon type, and an active matrix
of silicon pixels, each with a field effect transistor (FET), is
fabricated on a thin glass substrate. The panels are not
susceptible to radiation damage. As described above, in this dual
energy x-ray detection scheme, the soft x-ray image is captured by
the front phosphor/panel 18, and the hard x-rays are transmitted to
the rear phosphor/panel 22 while the inter-detector filter 20
attenuates any soft x-rays that may pass through the front
phosphor/panel 18. Since the optical photons generated by the soft
x-ray phosphor are captured by the front panel 18 and are not
channeled to another location for detection, it is not necessary
that the inter-detector filter 20 and the hard x-ray detector panel
22 be fiber optic plates. In addition, it is not necessary that the
soft and hard x-ray conversion screens be of the same type nor that
these screens generate optical photons of different colors as in
the embodiment of FIG. 2.
For definiteness in the following description, it will be assumed
the red phosphor and green scintillator dual-energy x-ray detection
scheme of FIG. 2. However, as indicated above, it is believed, from
modeling, that the signals from the amorphous silicon panels would
be comparable to those calculated for the red phosphor and green
scintillator scheme.
The coefficient of variation (CV) of the measurement of the BMD is
calculated using the usual expressions: ##EQU1##
where t.sub.B and t.sub.T are the bone mineral and tissue areal
densities, .sigma.(t.sub.B) is the variance in the BMD measurement,
N.sub.pixels is the number of detector pixels in the region of
interest, and N.sub.P and N.sub.S are the numbers of x-ray quanta
absorbed per pixel in the phosphor and scintillator, respectively.
The partial derivatives of N.sub.P and N.sub.S are taken at
constant bone mineral or tissue densities. J is the Jacobian of the
transformation from the densities (t.sub.B and t.sub.T) as a
function of absorbed quanta (N.sub.P and N.sub.S) to the absorbed
quanta as a function of densities. The latter quantities, the
absorbed quanta as a function of the densities, are more convenient
to calculate. It is assumed that .sigma..sup.2 (N.sub.P)=N.sub.P
and .sigma..sup.2 (N.sub.S)=N.sub.S.
The numbers of x-ray quanta (summed over energy) absorbed per pixel
in the phosphor and the scintillator (N.sub.P and N.sub.S,
respectively are approximately 600-900 for small BMD and decrease
to 400-700 for 1.5 g/cm.sup.2 BMD. The numbers of detected photons
per pixel (red from the phosphor 24 and green from the scintillator
28) are much larger.
The CV is quite sensitive to the partial derivatives of the numbers
of absorbed x-ray quanta per pixel that appear in the Jacobian. The
slopes of the absorbed quanta curves as functions of bone mineral
and tissue density have large negative values for small densities,
and this results in low CV values for small BMD.
The value of the BMD can be inferred from the number of detected
red and green photons from the phosphor and the scintillator and by
using the average attenuation coefficients of tissue and bone
mineral. Let .tau..sub.P and .tau..sub.S be the energy-averaged
attenuation coefficients of tissue weighted by the number of x-ray
quanta absorbed by the phosphor 24 and the scintillator 28,
respectively. Let .beta..sub.P and .beta..sub.S be the
energy-averaged attenuation coefficients of bone mineral weighted
by the number of x-ray quanta absorbed by the phosphor and the
scintillator, respectively. After passing through a tissue density
of t.sub.T and a BMD of t.sub.B, the x-rays that are absorbed by
the phosphor and the scintillator result in red and green photon
fluences of
where R.sub.O and G.sub.O are the red and green fluences in the
absence of the bone and tissue materials. The tissue density
t.sub.T can be eliminated, and the two equations can be solved for
the BMD,
where .rho.=.tau..sub.P /.tau..sub.S and .delta.=.rho..beta..sub.S
-.beta..sub.P. In effect, the soft tissue contribution has been
removed from the hard x-ray signal.
The inferred BMD is inaccurate at larger values of BMD, and this
indicates the need for a calibration phantom as discussed
below.
Contour plots of the number of x-rays absorbed by the phosphor 24
and the scintillator 28, in units of 10.sup.3 quanta per pixel,
plotted as functions of the tissue and bone mineral densities in
units of g/cm.sup.2, indicate that the x-ray quantum noise is
higher than the visible photon noise throughout the range of tissue
and bone densities.
For a 20 mm region of interest, the CV values that have been
plotted are less than 2% for the range of BMD that is typically
encountered in lumbar spine and proximal femur densitometry.
In order to improve the accuracy of the inferred BMD, a calibration
phantom was implemented in the computer model. Traditional
calibration phantoms are commonly composed of Plexiglas to simulate
soft tissue and aluminum to simulate bone mineral. As listed inn
Table 1, Plexiglas has composition of C.sub.5 H.sub.8 O.sub.2 and a
density 1.19 g/cm.sup.3. The x-ray attenuation of Plexiglas is
similar to that of soft tissue.
The atomic number (13) of aluminum is similar to the
weight-averaged atomic number (14.1) of hydroxyapatite (Ca.sub.5
P.sub.3 O.sub.13 H), the major constituent of bone mineral.
However, the attenuation coefficients of aluminum and
hydroxyapatite significantly differ at the soft and hard x-ray
energies typically used for dual-energy bone densitometry. A search
of readily available materials revealed that TiO.sub.2 is better
calibration phantom for hydroxyapatite than is aluminum. For an
appropriate thickness of TiO.sub.2, the attenuation coefficients at
the soft and hard x-ray energies are in good agreement with those
of hydroxyapatite.
The attenuation coefficients of tissue and bone mineral were
replaced by those of Plexiglas and TiO.sub.2 in the computer model
used above and the same procedure was used to calculate the x-rays
absorbed in the phosphor 24 and scintillator 28 and the resulting
red and green photons. The calibration quantities .rho. and .delta.
were derived for the calibration phantom.
The ratio of the .rho. values for the tissue/bone case and the
Plexiglas/TiO.sub.2 case and the corresponding ratio of the .delta.
values were used to correct the .rho. and .delta. values that were
used to calculate the BMD. This was done by fitting a smooth
surface, that was a polynominal function of the tissue and bone
mineral densities, to the surfaces. The errors in the inferred BMD
were less than 2% for the quadratic and cubic fits except for very
small values of bone mineral and tissue densities. Thus a
reasonably low-order polynominal correction can be made to the
inferred bone density by using the Plexiglas/TiO.sub.2 calibration
phantom.
Although the invention has been described above in relation to
preferred embodiments thereof, it will be understood by those
skilled in the art that variations and modifications can be
effected in these preferred embodiments without departing from the
scope and spirit of the invention.
* * * * *